Radiographic imaging devices (also known as radiographic imagers) are generally used to radiograph an object or to produce the image of a radiating source. They consequently find applications in non-destructive testing, radiography and tomography of objects or sources of radiation, gammagraphy, neutronography and in any imaging technique using X, beta, gamma, protons, neutrons, and the like. ionizing radiation. These devices can address numerous requirements running from the medical field to astronomy and to the inspection of packages for safe transport.
Hereinafter the term radiation (or ray) is used to refer to all ionizing radiation that can be used for imaging: X, beta, gamma, protons, neutrons, and the like. Similarly, the expression radiographic imaging will be used regardless of the type of radiation concerned.
Generally speaking, radiography of an object is effected by placing the object between a source of radiation and a radiographic imager. As for the imaging of a source of radiation, it is effected by placing a collimator between the source and the radiographic imager. The collimator can simply be a small hole and the image of the source is then formed on the imager, but this collimator can be more complex (for example in the case of multi-hole coded openings, slots, etc. or in the case of penumbra imaging) and digital processing of the images obtained can be necessary to reconstruct the image of the source.
Acquisition of radiographic images of the source (or the object) at various angles allows reconstruction in three dimensions of the source (or the object) by digital fusion of the data provided by the various views (this technique is called tomography). This is the case, for example, with single photon emission computed tomography (SPECT).
Radiographic imaging has grown considerably since the discovery of X rays. Initially, radiographic films were coupled with reinforcing screens the function of which was to convert the radiation into light to facilitate recording thereof by the film. Films have progressively been replaced by light-sensitive electronic sensors (such as CCD (charge-coupled devices), photodiodes, etc.) and the reinforcing screens have been replaced by scintillators. Optical coupling between the scintillator and the electronic sensor is generally effected by lenses or optical fibers with a reduction of image size and/or amplification of light where appropriate.
U.S. Pat. Nos. 6,031,892 and 6,285,739, for example, describe devices using a scintillator crystal, a CCD sensor and coupling optics.
In radiographic imaging devices of this type, the quality of the images depends mainly on the sensitivity and the spatial resolution of the combination of the scintillator, the coupling optics and the image sensor and on the level of the flux of photons producing the signal. The ionizing radiation deposits energy in the scintillator which releases it in the form of light detected by the electronic image sensor. The level of noise in the image (and therefore the signal to noise ratio) are thus mainly a function of the density of the scintillator, the effective section of the material from which it is produced (vis à vis the type and energy of the ionizing radiation concerned), the number of light photons emitted by the scintillator per unit dose deposited in the material by the radiation (luminous efficiency), the transparency of the scintillator to its own light radiation, the optical coupling between the scintillator and the sensor and the performance of the sensor itself.
For its part the spatial resolution is directly linked to the thickness of the scintillator. The spatial enlargement of the signal in the scintillator reflects the spreading of the deposits of energy of the incident rays within it (phenomena of diffusion of electrons and secondary photons created by interaction of the radiation with the scintillator). The thicker the scintillator, the greater the spreading of the deposit of energy. Increasing the thickness of the scintillator increases the sensitivity and the signal to noise ratio of the image at the same time as degrading spatial resolution. It is therefore difficult to produce a device that is both sensitive and of high resolution. The difficulty increases with highly penetrating radiation because it is necessary to increase the thickness of the scintillator to capture a sufficient proportion of the incident rays and to achieve good detection quantum efficiency (DQE).
Scintillators can be classified into two main families: organic or plastic scintillators (for example phosphor screens) and scintillator crystals: CsI, NaI, germanate of bismuth (usually called BGO), silicates of lutecium (LSO, LYSO), and the like.
Scintillator crystals have an advantage over organic scintillators because, in most cases, their density and their effective section of interaction with the radiation are greater. The thicknesses to be used to achieve a given detection efficiency are significantly reduced by this and the spatial resolution improved. Moreover, some crystals have properties that are very useful for radiographic imaging: very short light emission, allowing fast and repetitive recordings, very high luminous efficiency, allowing losses in the optical coupling. Crystals that perform well are described for example in French patent no. FR 2 874 021 and European Patent Publication No. EP 1 754 981.
At present the technology for manufacturing these crystals offering very good performance (high density and high luminous efficiency), such as BGO, LSO and LYSO, do not allow plates to be produced with sufficient dimensions for routine applications (the typical requirement is 300×300 m2, or even 400×400 mm2). LYSO and BGO, for example, can at present be obtained in plates having a uniform luminous efficiency and no major defects up to dimensions of approximately 60×60 mm2 and 80×160 mm2, respectively (and for thicknesses up to 30, or even 40 mm).
To produce large-format scintillators (exceeding 100×100 mm2) it is possible to use crystals of CsI or of NaI that are obtained with typical dimensions of 200×200 mm2 and 400×400 mm2, respectively. However, numerous defects remain present in crystals of CsI and NaI with these dimensions and these crystals are more or less hygroscopic. Moreover, the performance of the imaging device is less good with crystals of CsI or of NaI than with more dense crystals (such as BGO, LSO, LYSO) for the reasons stated above.
U.S. Pat. No. 7,057,187 and French patent No. FR 2 625 332 describe radiographic imaging systems using a scintillator crystal (preferably of CsI) optically coupled to a CCD sensor for X-rays with an energy from 1 to 100 keV. The crystal thicknesses are respectively 50 μm and 1 mm. The thickness of the scintillator has been limited to preserve good spatial resolution, to the detriment of sensitivity. This leads to a degraded quality radiographic image, to an increase in the exposure time necessary to obtain a sufficient detection statistic, and to an increase in the power of the source. This can be harmful in the case of medical applications, for example, in which the dose received by the patient must be reduced to that strictly necessary. This can also represent a penalty for inspecting packages because for the system to be operational the immobilization time is decisive.
The publication “A high quantum efficiency prototype video based portal imaging system” by Samant et al., published in Journal of X-ray Science and Technology 14 (2006, pp. 161-175) gives an example of a CsI scintillator 12 mm thick and with an area of 170×170 mm2 coupled to a CCD by means of a photographic lens. This device is limited by the available scintillator dimensions. Moreover, it is difficult to fabricate a large scintillator crystal without any imperfections in the bulk of the material. In the example cited, the crystal has approximately 20 to 30 imperfections (air bubbles) distributed randomly in the bulk of the scintillator. These defects diffuse the light generated by the scintillator and induce (unwanted) intensity peaks in the final image. However, to fabricate small and/or thin scintillators it is possible to select and to cut the crystal in portions free of defects.
Another known solution for producing large format scintillators is to use the segmentation technique, which allows good spatial resolution to be obtained at the same time as preserving sufficient sensitivity in high-energy radiographic imaging devices. This technique is routinely applied to scintillators of BGO, LSO, LYSO, BaF2, type, and the like.
Using this technique, the face of the scintillator receiving the radiation is divided into small elements 50 (see FIG. 9) that are optically separate and which therefore each transmit the light generated within them (photons, generally in the visible range). Each element thus forms a light guide and the elements are assembled mechanically to form the scintillator 52 (FIG. 8).
In practice, the small elements 50 are separated by opaque walls in order to prevent the light produced in each of them from reaching the adjacent elements. The light can be guided toward the exit face 54 of the scintillator by treatment of the lateral faces 56 of the scintillator elements (the faces adjacent the similar other elements, see FIG. 9). This treatment is generally adapted to produce specular reflection (furthermore entailing polishing of the lateral faces of the segments) so as to optimize the luminous efficiency. The light pickup is often in contact with the scintillator and the optical coupling is effected over a large numerical aperture (for example in the case of coupling the scintillator to a block of optical fibers or to a photomultiplier). However, in the case of optical coupling over a smaller numerical aperture (in the case of a system using a CCD and a photographic lens), it is more suitable (and less costly) to encourage diffusion from the lateral faces of the segments to increase the luminous efficiency in the direction of the axis of the segments (see for example the paper by Quillin and Aedy, “A pixelated BGO scintillator array for high energy flash radiography”, Nuclear Symposium Conference Record, IEEE 2004, Vol. 2, pp. 794-797.
The rear face 58 (opposite the light pickup) is for its part frequently treated to reflect light and increase the luminous flux directed toward the detector (this can be a metal deposit or simply white paint).
In some cases, lateral walls 60 are added between the scintillator elements 50 (see FIG. 10) to reduce the diffusion of secondary particles (for example electrons, photons). These walls generally consist of metal (aluminum, steel, tungsten).
The patents and patent applications U.S. Pat. No. 3,344,276, GB 2 167 279, GB 2 034,148, U.S. Pat. Nos. 5,773,829, 5,329,124, 6,344,649, US 2005/0104000, U.S. Pat. No. 7,238,945 describe segmented scintillators constructed in accordance with this principle.
The main drawback of segmented scintillator radiographic imagers is their cost, as they necessitate the laborious cutting and assembly of a very large number of elements as the pixel size must be close to the required spatial resolution. Moreover, it is difficult to obtain a perfectly regular assembly in which the sensitivity does not vary too much from one segment to another. Correcting these defects entails using subsequent digital image processing, of a more or less sophisticated nature, which induce a further loss of the quality of the final image (signal to noise ratio and effective resolution).
Finally, different structures have been proposed for optimizing the optical coupling between the scintillator and the sensor, the idea being to limit part of the spread of the light coming from the scintillator and to guide it toward the detector. U.S. Pat. Nos. 6,881,960 and 7,112,797 propose solutions with a one-piece scintillator semi-segmented over part of its thickness. U.S. Pat. No. 5,753,918 describes curved scintillator elements with reflecting or diffusing optical treatments. U.S. Pat. No. 6,737,103 proposes to add a matrix of optical microlenses to guide the light produced in the scintillator. These solutions increase the cost of the scintillator and do not solve the problem of large format crystalline scintillators. Moreover, as in segmented scintillators, this structuring leads to defects that are a problem in the radiographic images.